For diagnostic examination and for interventional procedures, for example, in cardiology, radiology, and also in surgery, X-ray systems are used for imaging. X-ray systems 16, as shown in FIG. 1, include an X-ray tube 18 and an X-ray detector 17, for example, arranged together on a C-arm 19, a high voltage generator to generate the tube voltage, an imaging system 21 (e.g., including at least one monitor 22), a system control unit 20, and a patient couch 23. Systems with two planes (e.g., two C-arms) are likewise used in interventional radiology. Flat panel X-ray detectors may be used as X-ray detectors in many areas of medical X-ray diagnostics and intervention (e.g., in radiography, interventional radiology, cardiac angiography), and also for imaging in therapy in the context of monitoring and planning of radiation or for mammography.
Present day flat panel X-ray detectors may be integrating detectors and may be based on scintillators that convert X-ray beams into comparatively low-energy radiation, for example, into visible light. This light is converted into an electrical charge in matrices of photodiodes. The matrices are then read line by line via active control elements. FIG. 2 shows the basic design in current use of an indirectly converting flat panel X-ray detector, including a scintillator 10, an active readout matrix 11 made of amorphous silicon with a plurality of pixel elements 12 (with a photodiode 13 and a switch element 14), and control and readout electronics 15 (see, for example, M. Spahn, “Flat detectors and their clinical applications,” Eur Radiol. (2005), 15: 1934-1947).
Depending on the quality of the beam, the quantum efficiency for a scintillator made of CsI with a layer density of, for example, 600 m is between 50% and 80% (see, for example, M. Spahn, “Flat detectors and their clinical applications,” Eur Radiol (2005), 15: 1934-1947). The local frequency dependent DQE(f) detective quantum efficiency has an upward limit as a result thereof and for certain pixel sizes of, for example, 150 to 200 □m, and, for the local frequencies of interest for practical applications of 1 to 2 lp/mm, falls well below this limit. In order to allow new applications (e.g., dual energy, material separation), but also to further increase the quantum efficiency, the potential of counting detectors or energy discriminating counting detectors mainly based on directly converting materials (e.g., CdTe or CdZnTe=CZT) and of ASICs with contacts (application specific integrated circuits such as a CMOS technology design) is increasingly used. Other materials such as Si or GaAs may likewise be of interest for specific applications.
An example of a design for such counting X-ray detectors is shown in FIG. 3. X-rays are converted in the direct converter 24 (e.g., CdTe or CZT), and the charge carrying pairs that are generated are separated via an electric field that is generated by a shared top electrode 26 and a pixel electrode 25. In one of the pixelated pixel electrodes 25 of the ASIC 27, the charge generates a charge pulse, the level of which corresponds to the energy of the X-ray quantum and that, if the charge pulse is above a defined threshold value, is recorded as a counting event. The threshold value serves the purpose of distinguishing an actual event from electronic noise or, for example, suppressing k fluorescence photons in order to avoid multiple counts. The ASIC 27, a corresponding section of the direct converter 24, and a coupling between the direct converter 24 and the ASIC 27 (e.g., in directly converting detectors using bump bonds 36) each form a detector module 35 with a plurality of pixel elements 12. The ASIC 27 is arranged on a substrate 37 and is connected to peripheral electronics 38. A detector module 35 may also include one or a plurality of ASICs 27 and one or a plurality of parts of a direct converter 24, selected as required.
FIG. 5 shows the general layout of a counting pixel element 12. The electrical charge is collected via the charge input 28 in the pixel element 12 where the electrical charge is amplified with the aid of a charge amplifier 29 and a feedback capacitor 40. In addition, on the output, the pulse shape may be adjusted in a pulse shaper (e.g., filter) (not shown). An event is then counted by moving a digital memory unit 33 (e.g., a meter or counter) up by one when the output signal is above a settable threshold value. This is established via a discriminator 31. In principle, the threshold value may also be provided in analog form, but is currently applied via a digital to analog converter (DAC) 32 and is thus variably adjustable within a certain range. The threshold value may either be adjustable locally on a pixel by pixel basis, as shown via the (local) discriminator 31 and the (local) DAC 32 or globally for a plurality of/all pixel elements via, for example, a global discriminator and DAC. The readout may subsequently ensue via a control and readout unit or peripheral electronics 38.
In addition to a global DAC that serves, for example, to adjust a specific keV threshold for an entire detector module or the entire X-ray detector, a further pixel by pixel adjustment that is intended to correct pixel to pixel fluctuations (e.g., fluctuations of amplifiers 29, local material non homogeneities in the detector material, etc.) may be necessary. This pixel by pixel calibration or correction DAC may have a considerably higher resolution than the global DAC and is adjustable, for example, across a keV range within which pixel to pixel fluctuations are expected (e.g., 6 keV). If such a calibration or correction DAC is provided, it is then advantageous to design the global DAC and the correction DAC separately due to the aforementioned different resolutions. The global DAC may then be applied with a rather lower resolution (e.g., 2 keV/bit) that generates a voltage that is applied on each pixel element of the detector module or for all the detector modules in a detector and on which a pixel by pixel corrected voltage is superimposed pixel by pixel via a higher resolution correction DAC (e.g., 0.1 keV/bit or 0.5 keV/bit). If a plurality of threshold values and counters are provided per pixel element (e.g., spectral imaging), then a plurality of global DACs are necessary. It may be advantageous to provide a calibration or correction DAC for each discriminator in case, for example, the circuit works in a non linear manner.
FIG. 6 shows a diagram for an entire array of counting pixel elements 12 (e.g., 100×100 pixel elements each of 180 μm). In this example, the size of the array would be 1.8×1.8 cm2. For large scale X-ray detectors (e.g., 20×30 cm2), a plurality of detector modules 35 are combined (e.g., around 11×17 modules would produce this area) and connected via shared peripheral electronics. For the connection between the ASIC and the peripheral electronics, through silicon via (TSV) technology, for example, is used to provide quadrilateral mounting of the modules side by side.
In the case of counting and energy discriminating X-ray detectors, two or more (e.g., four, as shown in FIG. 7), different threshold values are inserted per pixel by four pairs including a DAC 32 and discriminator 31. The level of the charge pulse, corresponding with the pre-defined threshold values (e.g., discriminator threshold values), is classified into one or a plurality of the digital memory units 33 (e.g., counters). The X-ray quanta counted in a specific energy range may then be obtained by subtracting the counter contents from two corresponding counters. The discriminators 31 may be adjusted, for example, with the aid of digital to analog converters for the entire detector module or pixel by pixel within set limits or ranges. The counter contents of the pixel elements 12 are read out in succession in a modular manner via a corresponding readout unit.
An increase in the spectral resolution by adding further threshold values through additional discriminators and a corresponding DAC or memory unit is accompanied by increased space requirements on the ASIC, such that a random energy discrimination may not be possible at the present time for reasons of space.
It would be possible to reduce the size of the structures that are arranged on the ASIC, for example, by moving from the 180 nm technology to a 130 nm or 90 nm technology or even smaller. As a result thereof, the space requirement for the electronics components on the ASIC would be reduced to implement an energy threshold, so that more energy thresholds altogether may be achieved on the ASIC. However, this procedure would provide a large technological advance, which with only low production runs, may be cost intensive and therefore unprofitable. In medical imaging, the reduction in the sizes of the structures on the ASIC would not result in any reduction in the size of the ASIC itself or of the production costs since the detector area and hence the size of the ASIC are to be retained.